An Electrosurgical System

ABSTRACT

An electrosurgical system has an electrosurgical generator and a bipolar electrosurgical instrument, the generator being arranged to perform a treatment cycle in which radio frequency energy is delivered to the instrument as an amplitude-modulated radio frequency power signal in the form of a succession of pulses characterized by successive pulses of progressively increasing pulse width and progressively decreasing pulse amplitude. There are periods of at least 100 milliseconds between successive pulses, and the treatment cycle begins with a predetermined pulse mark-to-space ratio. Energy delivery between pulses is substantially zero. Each burst is of sufficiently high power to form vapor bubbles within tissue being treated and the time between successive pulses is sufficiently long to permit condensation of the vapor.

FIELD OF THE INVENTION

[0001] This invention relates to an electrosurgical instrument, to anelectrosurgical method and to an electrosurgical system. Morespecifically, the present invention provides a bipolar radiofrequency(r.f.) electrosurgical output, the characteristics of which producefaster, more controlled and more effective sealing of vascularstructures such as might be performed during endoscopic surgery, and thelike.

BACKGROUND OF THE INVENTION

[0002] The use of r.f. current to effect the cutting and coagulation ofbody tissues has been known for many years, and comes under the broaddescription of electrosurgery. Two techniques to deliver the r.f.current to the tissues are in common usage today.

[0003] The first of these, monopolar electrosurgery, involves the use ofan active (tissue treatment) electrode and a remote return (or neutral)electrode (or pad) placed on an external surface of the patient's body.Current flows from the active electrode, through the target site, andthrough any other tissue lying in the path between the target site andthe return electrode. This arrangement introduces the potential foroff-site burns, in other words tissue burns occurring at sites otherthan the target site. The medical literature includes references tonumerous instances of capacitive coupling of the r.f. current to otherinstruments causing burns, direct coupling to tissue due to insulationfailure, burns along the current path through the patient's body, andthose occurring at the site of application of the return pad.

[0004] The second technique is known as bipolar electrosurgery. Bipolarelectrosurgery involves the containment of current flow local to atarget site by incorporating both the active and return electrodes closetogether, typically at the tip of the surgical instrument. Thisarrangement avoids the need for current flowing through the body tocomplete the electrical circuit, and hence eliminates the risks ofoff-site burns. The use of bipolar electrosurgery is, therefore,preferred where safety is of greatest concern, particularly whenapplying r.f. current close to vital structures, or when visualisationis limited such as during endoscopic surgery. As a result, bipolarcoagulation or sealing of vessels during endoscopic surgery has become acost-effective and easy to use alternative to the mechanical sealing ofblood vessels using metal clips, staples or ligatures.

[0005] Normally, the electrosurgical instrument used for bipolarcoagulation consists of a pair of forceps, in which each jaw of theforceps is an r.f. electrode. Depending on the size of the forceps, andhence the amount of tissue included in the circuit, the applied powercan typically vary between 1W and 50W. The most significant problemsencountered, when using conventional bipolar r.f. electrosurgery, arerelated to the distribution of energy throughout the tissue graspedbetween the forceps. As a result of these limitations, surgeons willcommonly apply r.f. energy well beyond that necessary for effectivelysealing a blood vessel, in theory to ensure complete sealing and toreduce the risk of bleeding when the vessel is subsequently divided.This leads to an excessive spread of the coagulation to adjacenttissues, and increases the risk of the forceps jaws becoming stuck tothe tissue. This sticking can be sufficiently severe to cause coagulatedtissue to be torn away when releasing the forceps, leading to damage ofuntreated areas of the vessel, and significant bleeding.

[0006] The industry standard for the coagulation output of a bipolarr.f. electrosurgery generator is a maximum power in the region of50W-70W, with a specified load curve between 10 ohms and 1000 ohms. Thispower is normally delivered as a continuous, low crest factor waveformsuch as a sine wave. Peak voltages from such a generator can be as highas 1000V peak-to-peak. It has now been recognised, however, that lowervoltages reduce the propensity to stick or carbonise the tissue whencoagulating. Maximum voltages of up to 400V peak-to-peak are now moreusually used in modern designs. The low impedance matching capability ofthis type of generator is limited, with maximum current deliverytypically being in the region of 1.5A at full power.

[0007] Despite these advances, none of the known bipolar r.f. generatorsovercomes the problems of differential energy absorption within thetissue due to the variation in tissue impedance, the geometry of theforceps jaws, the presence of conductive fluids and tissue compression.As a result, coagulation is inevitably taken to the desiccation point,at which the tissue becomes dried out as fluids are boiled off, with anattendant elevation in the temperature of the forceps jaws. The cause oftissue sticking is the elevation in electrode temperature above 70-80°C. As this is more likely to occur because of the variables encounteredduring use, it is particularly likely to occur when the vessel to betreated is contained within the high impedance of a fatty layer, as iscommonly encountered in vascular pedicles. The fatty layer effectivelyinsulates the lower impedance vascular structure, so that incompletesealing and excessive application are both more likely to occur.

[0008] For these reasons, it would be desirable to deliver bipolar r.f.electrosurgical energy in an improved way for coagulating tissues. Itwould be particularly desirable to provide more controlled absorption ofenergy throughout the tissue to be treated, largely irrespective ofvariables encountered during use, so that the problems of incompletevessel sealing within fatty pedicles, tissue sticking and excessivethermal margin can be overcome. It would further be desirable to providean improved bipolar r.f. electrosurgical output through an instrumentsuch as that disclosed in U.S. Pat. No. 5,445,638 during endoscopicsurgery.

[0009] Electrosurgical instruments have been proposed to resolve theproblems of sticking. U.S. Pat. Nos. 3,685,518, 4,492,231 and 6,059,783all describe methods of heat removal by constructing the electrodes ofsufficient thermal capacity, and/or by the use of thermally-conductivematerials to dissipate heat. U.S. Pat. No. 5,885,281 describes the useof coatings to minimise the effects of sticking.

[0010] Impedance and temperature-based r.f. generator control isdescribed in U.S. Pat. No. 5,496,312. Our U.S. Pat. No. 5,423,810describes an impedance-controlled, bipolar cauterising output based onvariations in the oscillator carrier frequency according to tissueimpedance.

[0011] U.S. Pat. No. 6,033,399 discloses an electrosurgical generatorcapable of applying output power to surgical graspers in a manner suchthat the power level varies cyclically between low and high values inresponse to the changing impedance of the grasped tissue being treated,until the tissue is fully desiccated.

[0012] These techniques have had moderate success in terms of preventingsticking. One method of counteracting the negative temperaturecoefficient of resistance (NTCR) effect which tissue exhibits duringcoagulation is to introduce a positive temperature coefficient ofresistance (PTCR) material, which new material is dominant. PTCRmaterial produces the opposite effect to current hogging so that,instead of current hogging, the predominant effect would then be one ofcurrent sharing. Whilst it might be possible to coat the electrodes witha PTCR material, the material would dissipate heat and heat up theelectrodes. Alternatively, a dielectric layer could be introduced with apositive temperature coefficient of impedance. This has the attractionof little or no heat dissipation, but unfortunately is very difficult torealise due to the lack of suitable materials.

SUMMARY OF THE INVENTION

[0013] According to a first aspect of the invention, an electrosurgicalgenerator comprises a source of r.f. energy, at least a pair of outputterminals for connection to a bipolar electrosurgical instrument and fordelivering r.f. energy from the source to the instrument, and a pulsingcircuit for the source, wherein the pulsing circuit and the source areso arranged as to deliver into a resistive load, when connected acrossthe output terminals, an amplitude-modulated r.f. signal at the outputterminals in the form of a succession of pulses which is characterisedby successive pulses being of decreasing power amplitude and ofincreasing pulse width. Preferably, the increase in the pulse width ofsuccessive pulses is related to the decrease in amplitude. This may besuch that the energy delivered in each pulse is substantially fixed frompulse to pulse over at least a major part of the succession of pulses.The decrease in pulse amplitude and the increase in pulse width may beeffected by control circuitry which includes a sensor circuit formonitoring at least one electrical parameter associated with the outputterminals, so that the pulse amplitude and pulse width are controlledaccording to the monitored electrical parameter. Preferred electricalparameters in this respect include the load resistance and, in the casewhere the r.f. voltage applied to the load is limited to a predeterminedmaximum value, the time from the start of a respective pulse to theinstant at which the peak voltage reaches the predetermined maximumvalue. Generally, the maximum value is below 200V peak and is preferablyin the range of from 50V peak to 200V peak.

[0014] Typically, the pulse repetition rate is less than or equal to 5Hz and is preferably less than 1 Hz, the r.f. source and the pulsingcircuit being arranged to generate a succession of treatment pulses ofr.f. energy at the output terminals with the periods between successivesuch pulses being 100 ms, preferably 300 ms or longer. Between thepulses the delivered power is preferably substantially zero.

[0015] In the case of the pulse repetition rate being less than 1 Hz,the pulsing circuit and the r.f. source are arranged to generate asuccession of treatment pulses of r.f. energy at the output terminals,the periods between successive such pulses being 1 second or longer.

[0016] Typically, each pulse of the succession of pulses delivers atleast 2 joules of r.f. energy when the load resistance is in the rangeof from 10 Ω to 1 kΩ.

[0017] The treatment cycle may commence with a succession of pulses, asdescribed above, followed by a period of continuous energy delivery,which may be at least 1 second in duration.

[0018] Additional preferred features are contained in the accompanyingdependent claims.

[0019] According to a further aspect of the invention, anelectrosurgical generator comprises a source of r.f. energy, at least apair of output terminals for connection to a bipolar electrosurgicalinstrument and for delivering r.f. energy from the source to theinstrument, and a pulsing circuit for the source, wherein the pulsingcircuit and the source are so arranged as to deliver into a resistiveload, when connected across the output terminals, an amplitude andmodulated r.f. power signal in the form of a succession of pulses havinga pulse duty cycle which increase progressively. Progressive increase isaccompanied by progressive decrease in peak power. Indeed, the increasein peak power may be such that the average delivered power over atreatment cycle remains substantially constant. The generator preferablyincludes a sensing circuit associated with the output terminals forsensing the above-mentioned electrical parameter. The pulsing circuitmay then be arranged automatically to adjust the duty cycle of thesignal pulses in response to the sensed electrical parameter. Thesensing circuit may, alternatively, or additionally, be arranged to beresponsive to an identification element housed in an instrumentconnected to the output terminals. It may be arranged to detect aninitial value of load impedance associated with the start of thetreatment cycle, the pulsing circuit preferably being arranged such thatthe initial pulse duty cycle increases with increasing sensed initialload impedance.

[0020] The invention also includes a method of electrosurgicallycoagulating tissue between the electrodes of a bipolar electrosurgicalinstrument, in which r.f. energy are applied to the tissue via theelectrodes in a succession of pulse bursts of progressively increasinglength and progressively decreasing amplitude. Typically, the increasein the length of successive bursts is related to the decrease inamplitude of those bursts.

[0021] According to another aspect of the invention, there is providedan electrosurgical system comprising the combination of anelectrosurgical generator as described above, together with a bipolarelectrosurgical instrument coupled to an output of the generator. Theinstrument may be a pair of forceps. In addition, the instrument ispreferably removably connectible to the generator and includes aninstrument identification element, as mentioned above. The system maycomprise the generator together with a plurality of bipolarelectrosurgical forceps instruments which are selectively connectible tothe generator, each containing respective different identificationelements. The electrodes of the instruments define different tissuecontact areas, the respective identification elements being arranged toco-operate with the sensing circuit and/or the pulsing circuit so as toset the pulse width or pulse duty cycle to a lower value for aninstrument with electrodes defining a comparatively large tissue contactarea and to a higher value for an instrument with electrodes defining acomparatively small tissue contact area. Additionally or alternative,the sensing circuit and/or the pulsing circuit are selected andconfigured in combination with the identification elements to decreasethe pulse frequency when an instrument with a comparatively large tissuecontact area is selected.

[0022] According to yet another aspect of the invention, anelectrosurgical generator comprises a source of r.f. energy, at least apair of output terminals for connection to a bipolar electrosurgicalinstrument and for delivering r.f. energy from the source to theinstrument, and a pulsing circuit for the source, wherein the pulsingcircuit and the source are arranged to deliver into a resistive loadacross the output terminals an amplitude-modulated r.f. signal at theoutput terminals in the form of a succession of pulses characterised bythe periods between successive pulses in the signal being at least 100ms and by a predetermined mark-to-space ratio Preferably, the depth ofamplitude modulation is substantially 100%, with a pulse mark-to-spaceratio of less than 1:1.

[0023] When a resistive load is coupled across the output terminals ofthe generator, the r.f. current during each of a number of successivepulses may reach at least 3 amps r.m.s.

[0024] In one preferred generator in accordance with the invention, thecircuitry is arranged such that the peak voltage of theamplitude-modulated r.f. signal remains below 200 volts when a resistiveload is connected across the output terminals, the r.f. energy deliveredin each pulse being at least 2 joules when the resistive load is in therange of from 10Ω to 1 kΩ.

[0025] According to an additional aspect of the invention, anelectrosurgical generator comprises a source of radio frequency (r.f.)energy, at least a pair of output terminals for connection to a bipolarelectrosurgical instrument and for delivering r.f. energy to theinstrument, a pulsing circuit for the source, and control circuitryincluding means for monitoring at least one electrical parameterassociated with the output terminals, wherein the arrangement of thecontrol circuitry, the pulsing circuit and the source is such that, withthe output terminals connected to a resistive load, the controlcircuitry causes the source to deliver into the load anamplitude-modulated r.f. power signal which, at least in an initialperiod, is a succession of pulses with a predetermined initial pulseduty cycle and, in a subsequent period, has a different characteristic,the transition from the initial period to the subsequent period beingcontrolled by the control circuitry in response to the at least onemonitored parameter. The control circuitry may be arranged to cause ther.f. power signal, during the subsequent period, to provide continuousenergy delivery or more nearly continuous energy delivery than duringthe initial period but, more commonly, the r.f. power signal isdelivered as an amplitude-modulated signal which, during at least partof the above-mentioned subsequent period, has a second predeterminedpulse duty cycle which is greater than the initial pulse duty cycle.Generally, the peak power during the subsequent period is less thanduring the initial period. In one particular preferred embodiment, thepulse duty cycle is fixed at a first predetermined pulse duty cycleduring the initial period and at a second, greater predetermined pulseduty cycle during the subsequent period, the subsequent period followingthe initial period directly. As an alternative, the pulse duty cycle ofthe r.f. power signal may increase in more than one step so that, forinstance, the signal starts with a low predetermined and fixed pulseduty cycle, then is switched to a pulse duty cycle which is greater thanthe first pulse duty cycle and with lower peak power and, subsequently,to a yet higher pulse duty cycle and yet lower peak power. As a furtheralternative, the pulse duty cycle may increase progressively,accompanied by progressively reducing peak power.

[0026] In this case, whether the treatment cycle performed using ther.f. power signal is a pulse signal followed by a continuous-wave (c.w.)signal, or a signal in which the pulse duty cycle is increased stepwiseor progressively, the peak power may be correspondingly reduced suchthat the average delivered power remains approximately constant over themajority of the treatment cycle, the cycle commencing with the initialperiod and ending when the r.f. power signal is terminated.

[0027] The transition from the initial period to the subsequent periodmay be controlled in response to a feedback signal representative ofenergy delivered into a resistive load, or one which is representativeof the resistance or impedance of the load. A feedback signal may beobtained by sensing the output voltage (peak voltage or r.m.s. voltage),the transition being controlled in response to a sensing signal from asensing circuit indicative of the output voltage exceeding apredetermined value, for instance. The predetermined value may be in theregion of 150V to 250V peak.

[0028] In the case of the generator having a switched mode power supplyoperating at a power supply switching frequency, the output voltagesensing circuit may be coupled to the power supply in such a way thatwhen the output voltage exceeds a predetermined value, pulsing of thepower supply is halted. The output voltage may then be sensed bymonitoring the driving pulses of the power supply, e.g. by counting thepulses. The counting output may then be used to control the pulse dutycycle and/or peak power of the r.f. power signal.

[0029] According to yet an additional aspect of the invention, a methodof electrosurgically coagulating tissue between the electrodes of abipolar electrosurgical instrument comprises the application of r.f.energy to the tissue via the electrodes in a succession of pulse burstswith a duty cycle of 40% or less, wherein the instantaneous r.f. currentat the start of each successive burst is higher than the instantaneousr.f. current at the end of the previous burst.

[0030] A further aspect of the invention provides an electrosurgicalsystem comprising an electrosurgical generator and a bipolarelectrosurgical instrument coupled to an output of the generator, thegenerator being such as to provide a succession of controlled bursts ofelectrosurgical energy to the instrument at a predetermined pulsemark-to-space ratio, wherein each burst has a sufficiently high power toform at least one vapour bubble within tissue being treated by theinstrument, and the time duration between successive bursts issufficiently long to permit recondensation of the or each vapour bubble,the peak delivered power being between the bursts being substantiallyzero. The time delay duration is generally at least 100 milliseconds andthe generator preferably has other features already mentioned.

[0031] The invention also includes an integrated electrosurgicalgenerator and instrument system, wherein the generator has the featuresreferred to above and the instrument comprises a pair of forceps.

[0032] According to yet a further aspect of the invention, there isprovided an integrated electrosurgical generator and instrument system,when the instrument is removably connectible to the generator andincludes an instrument identification element. The generator may haveany of the above-mentioned generator features and includes a sensingcircuit for sensing the identification element, the pulsing circuit ofthe generator being arranged automatically to adjust the mark-to-spaceratio of the signal pulses in response to the identification element assensed by the sensing circuit. The system may include a plurality ofbipolar forceps instruments which are selectively connectible to thegenerator and contain respective identification elements. Theinstruments have different tissue contact areas (defined by theinstrument electrodes) and the identification elements are selected suchthat, in combination with the sensing circuit and/or the pulsing circuitof the generator, the mark-to-space ratio is set to a lower value for aninstrument with electrodes defining a comparatively large tissue contactarea and to a higher value for an instrument with electrodes defining acomparatively small tissue contact area. The identification elements,the sensing circuity and the pulsing circuit are preferably selected andconfigured to decrease the pulse frequency when an instrument with acomparatively large tissue contact area is selected.

[0033] Yet another aspect of the invention provides a method ofelectrosurgically coagulating tissue between the electrodes of a bipolarelectrosurgical instrument in which controlled bursts of r.f. energy areapplied across the electrodes, each burst being of sufficiently highpower to form at least one vapour bubble within the tissue, and the timeduration between successive bursts is sufficiently long to permitrecondensation of the or each bubble.

[0034] The tissue behaves as a positive temperature coefficient ofresistance (PTCR) material by the application of r.f. energy at highpower across the electrodes of a bipolar instrument. The PTCR effect isproduced by exploiting the tendency for “current hogging” whereby, dueto a negative temperature coefficient resistance (NTCR), the applicationof r.f. energy to a region of tissue causes local temperature increaseswhich, in turn, causes localisation of current density, the r.f. currenttending to be concentrated at the areas of highest temperature,especially when, for instance, a thin section of tissue is graspedbetween electrodes formed as a pair of forceps. The PTCR effect isachieved by delivering sufficient power to the tissue that a vapourbubble is formed which, providing the applied voltage is substantiallybelow 300 volts peak, is substantially an electrical insulator. Since,now, the r.f. current must find paths around the vapour bubble, thematerial as a whole has exhibited a rise in impedance, effectivelygiving a PTCR characteristic. The dissipation of energy is thus moreevenly distributed and thermal coagulation occurs throughout the targettissue as a result.

[0035] A notable feature is that the highest temperatures, induced bythe highest current densities, occur within rather than on the surfaceof the tissue between the instrument electrodes.

[0036] Once vapour is formed, the highest current densities occur aroundthe edges of the vapour bubbles, causing further heating and expansionof the vapour bubble until the end of the respect pulse burst, asexpansion of the bubbles being such that the areas of highest currentdensity are forced into untreated regions of tissue below the tissuesurfaces. This reduces the risk of localised heating of the forceps jawsand hence reduces the risk of tissue sticking.

[0037] These effects result in preferential and more uniformdistribution of energy dissipation within the target tissue to provide amethod of treating tissue whereby a lateral margin of thermal effect isreduced and further that the coagulating effect on blood vessels can beobtained throughout other support tissues such as fatty connectivetissue. A further resulting advantage is that the surgeon is providedwith a more repeatable end-point of coagulation treatment despite thevariable conditions which may be encountered.

[0038] The control of the tissue effect may be obtained by altering thepulse characteristics depending on the specific instruments connected tothe generator with the effect of reducing the variables encounteredduring use. It is also possible to reduce the variables in the case of aforceps instrument embodiment by controlling the closure force exertedon the tissue.

[0039] In this connection and as described, the generator pulsingcircuit may be arranged automatically to adjust the mark-to-space ratioof the signal pulses in response to a sensing circuit associated withthe output terminals. The sensing circuit may be arranged to beresponsive to an identification element, such as an element having aparticular impedance, housed in an instrument connected to the outputterminal. Alternatively, the sensing circuit may be arranged to detectan initial value of a load impedance between the output terminals, whichvalue is associated with the start of r.f. energy application, the pulsecharacteristics being set according to the initial load impedance valuefor the duration of a treatment operation comprising a succession of thepulses. Typically, the pulsing circuit is arranged such that the initialpulse mark-to-space ratio increases with increasing sensed initial loadimpedance. In addition, the pulsing circuit may be arranged to adjustpeak power in response to the sensing circuit, the set peak powerdecreasing as the sensed initial load impedance increases. The pulsefrequency may also be adjusted by the pulsing circuit in response to thesensing circuit, the pulse frequency being increased with increasingsensed initial load impedance.

[0040] In the case of the instrument (which can include the connectingcable and its connector) containing an identification element such as acapacitor, resistor, or other coding element, the mark-to-space ratiomay be set according to the tissue contact areas of the electrodes, suchthat instruments with larger tissue contact areas cause the generator tobe set with a comparatively low mark-to-space ratio.

[0041] The invention will be described below in greater detail, by wayof example, with reference to the drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

[0042] In the drawings:

[0043]FIG. 1 is a graph illustrating the ideal behaviour of tissueimpedance against time during the application of bipolar r.f. energy;

[0044]FIG. 2 is a graph illustrating the compound behaviour of tissueimpedance against time as a result of the phenomenon of current hogging;

[0045]FIG. 3 is a schematic circuit diagram illustrating the currentdistribution density associated with current hogging when an r.f. sourceis applied across a laminar section of tissue;

[0046] FIGS. 4A-4D are schematic diagrams illustrating variations incurrent density when a vapour bubble is formed within a laminar sectionof tissue;

[0047] FIGS. 5-11 contrast the effect obtained on a tissue pedicle usingforceps operated conventionally and as part of a system in accordancewith the invention; and

[0048]FIGS. 12 and 13 are graphs illustrating the comparative efficiencyof energy delivery using forceps operated conventionally and as part ofa first system in accordance with the invention;

[0049]FIG. 14 is a diagrammatic representation of a system in accordancewith the invention, comprising an electrosurgical generator and aninstrument in the form of a pair of forceps;

[0050]FIG. 15 is a graph showing the average output power of anelectrosurgical generator as a function of load resistance, whenoperated in a continuous mode and in a pulsed mode with a 15% pulse dutycycle;

[0051]FIG. 16 is a graph showing the variation of pulse duty cycle andpeak power according to initial load impedance in one embodiment ofgenerator in accordance with the invention;

[0052]FIG. 17 is a graph showing the variation of pulse frequency withinitial load impedance in the same generator;

[0053]FIGS. 18 and 19 are graphs showing the variation of deliveredpower with time in second and third embodiments of the invention, and

[0054]FIG. 20 is a graph showing the variation of delivered power withtime in a fourth preferred embodiment of the invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT OF THE INVENTION

[0055] Referring to the drawings, FIG. 1 is a graph showing the idealbehaviour of tissue impedance against time during the application ofbipolar r.f. energy. The impedance is seen to fall during the initialphase of application as a result of heating of electrolytes in thevicinity of the tissue being treated. A minimum M is reached, followingwhich the impedance begins to rise as the tissue is desiccated andbecomes less conductive. Treatment, in terms of coagulation of thetissue, optimally occurs around the point M of minimum impedance.Continued delivery of energy beyond this point M merely serves toincrease the lateral margin, to increase the temperature of theapplication electrodes, typically a pair of forceps jaws, due toincreased steam generation and to increase the risk of tissue sticking.Increased ion mobility can cause a 60% impedance reduction over atypical temperature change of 37° C. to 100° C. In practice, however, a60% reduction is never seen since the tissue is never at a uniformtemperature.

[0056]FIG. 2 is a graph showing two solid line relationships whichillustrate how the tissue impedance may change at different pointsacross the contact areas of a typical bipolar forceps. Plot 1 isindicative of a point at which the impedance across the forcepsdecreases rapidly on application of power, such as that which may occurdue to the forceps jaws being closer together at one point along theirlength. As a result, this point of the contact area will take marginallymore power from the common bipolar r.f. power source. This, in turn,will cause heating at this point, with a further lowering of theimpedance, and a consequential increase in the power delivered at thispoint at the cost of other, higher impedance points of contact such asfor that point shown in Plot 2. This is the phenomenon known as currenthogging, and it is a feature of materials, such as conductive tissuefluids, which exhibit a negative temperature coefficient of resistance(NTCR). These individual characteristics will, of course, not be seen bythe common energy source, which will only see the combined effect of thetwo as indicated by the dotted line.

[0057] The first notable feature of the combined effect is that theimpedance minimum M¹ is less pronounced. The second notable feature isthat, when desiccation occurs, the positive rise in impedance withapplied power results in the opposite of current hogging, this beingknown as current sharing. This current sharing results in a convergenceof the two plots when desiccation starts to occur. It is for this reasonthat end-point determination of treatment can only be reliably detectedthroughout the tissue pedicle once the tissue reaches the point ofdesiccation, with the attendant unnecessary margin of effect, ahardening of the tissue, and electrode/tissue sticking.

[0058] The current hogging phenomenon can be more easily understood byconsidering two infinitely-small pieces of tissue to which the samepower source is applied, i.e. two pairs of electrodes connected inparallel to the same power source and applied to these two microscopicpieces of tissue. If one of these pieces of tissue has a marginallylower impedance than the other, it will take marginally more power.However, this marginal power increase in the lower impedance piece willresult in greater heating. Greater heating, as explained above, willresult in lower impedance. Thus, the power differential between the twopieces will increase, resulting in an even greater power differential.This is the current hogging phenomenon, and it always happens inmaterials with a negative temperature coefficient of impedance which, inthis instance, constitutes the electrolyte within the tissue. Apractical electrode applied to tissue will effectively have an infinitenumber of tissue sections behaving between these two extremes. Asalready stated, the electrical characteristics of each of these sectionswill have a tendency to converge at desiccation. The safest approach is,therefore, to use the point of desiccation as the end point for appliedpower, and this is easily detectable due to rapidly rising voltage atthe output of the generator, or by the lack of activity at the targettissue. However, this gives rise to the four problems mentioned earlier.The surgeon is, therefore, faced with the dilemma of trying to ensuretreatment is sufficient to seal vessels, versus the risk of tissuesticking and increasing the lateral thermal margin.

[0059] As already stated, the variation over the forceps surfaces is dueto initial impedance, temperature, electrical conductivity, tissuethickness and electrode surface area. Most of these variables are highlyinteractive and, therefore, difficult to isolate. The net effect,however, is one of current hogging and differential energy absorptionthroughout the tissue included within the grasp of the forceps. This isquite clearly illustrated when using forceps such as those described inthe above-mentioned U.S. Pat. No. 5,445,638. The region of coagulationcan be seen to start at one end of the forceps and to work its wayalong. This usually occurs due to the jaws not being parallel when theyare closed, such that the coagulation commences in the region of lowestimpedance (or closest proximity of the jaws) which will then exhibit thecurrent hogging phenomenon. There is, therefore, a decreased possibilityof viable coagulation along the full length of the jaws, withoutsticking occurring at the point at which coagulation commenced.

[0060] Practical trials show that the thicker the tissue treated, theless the propensity for current hogging. Current hogging occurs due toexclusive current paths. Consider an extreme case of amicroscopically-thin layer 1 of tissue something like a postage stamp,with an electrode 2 (shown only schematically) applied to the glue sideof the “stamp”, such as is illustrated in FIG. 3. If current is passedfrom an r.f. source 3 through a single point 4 at one corner to theother side, the current will preferentially go directly across the“stamp” to the electrode 2 on the other side. More significantly, nocurrent will travel across the “stamp” in any other region. Thus, anexclusive current path 4 is set up in the tissue. Thin tissue sections,therefore, dramatically increase the propensity for current hogging.

[0061] Preventing local temperature rises can reduce the effects ofcurrent hogging. As explained earlier, current hogging occurs due to thecyclical cause-and-effect of reduced impedance creating greater heat,causing reduced impedance. Spreading of heat over the contact surfaceswill reduce this cyclical event. The heat provided by a low impedancepoint, if spread, would reduce the impedance of adjacent points; and,therefore, decrease the possibility of current hogging. Using anelectrode surface that is highly thermally-conductive can do this, as istaught in the prior art.

[0062] Still more attractive is the removal of heat at thetissue/forceps contact surface to prevent the formation of hot spots, sothat tissue at the tissue/forceps boundary is kept at a lowertemperature, and tissue fluids are prevented from boiling. This measureensures that maximum temperature rises occur within the tissue pediclerather than at the surface, resulting in desiccation being limited towithin the tissue. Providing forceps jaws with a sufficiently massivethermal heat capacity can achieve this, as is also taught by the priorart.

[0063] The fundamental cause of tissue sticking is build-up of heatwithin the electrodes or forceps jaws. When an electrode reachestemperatures in excess of 80° C., sticking invariably happens, and isworsened when the tissue approaches desiccation. Power delivery aftercoagulation generates steam that quickly heats up electrodes. Theelectrodes are exposed to more than three times the energy dissipationto reach desiccation than they are to reach the pure coagulation point(shown as the minimum M in FIG. 1). Electrodes are, therefore, far morelikely to reach sticking temperatures when tissue is treated to adesiccation state.

[0064] The electrode-to-tissue interface is the energy transfermechanism to the target tissue. Given a fixed contact area, theelectrodes heat up if the electrode-to-tissue contact is in any wayelectrically resistive, and as a result of thermal conduction from thetreated tissue.

[0065] In tests using stainless steel or gold electrodes, tissue contactimpedance is of the order of 30% lower for gold than for stainlesssteel. This difference is attributed to the existence of oxide layers onthe steel electrode surface. The significance and potential benefit isunknown. This drop would, however, reduce the power dissipation at thispoint by a corresponding 30%. This is also taught by the prior art, inparticular in U.S. Pat. No. 5,885,281.

[0066] Obviously, the tissue next to the electrode surface will get hot.Thermal conduction from tissue to the treatment electrodes is dependenton temperature difference and time. The significant factor here is, ifthe entire volume of treated tissue is in thermal contact with theelectrodes, then a much greater proportion of the applied energy is usedto heat the electrodes.

[0067] As the treatment tissue thickness decreases, a greater proportionof the applied energy causes electrode heating, due to the shorterthermal conduction paths. However, as ever thinner tissue requires lesspower due to less volume, the two effects tend to cancel one anotherout, so that electrode temperature as a result of tissue thickness isrelatively constant. This, however, makes the assumption that tissueheating is performed uniformly throughout the tissue. In practice, thisthin layer of tissue will be particularly susceptible to the occurrenceof current hogging and formation of hot spots, due the greatervariations in the impedance between the electrodes or forceps jaws. Theissue then becomes one of local temperature rises, rather than bulktemperature rises of the electrodes.

[0068] Typical bipolar instrumentation designed for endoscopic use isinvariably limited in design, due to confines of the access ports.Standard entry port sizes of 5, 7.5 and 10 mm exist. The mechanicalaspects of designing such instruments invariably result in hingeddesigns with a long length to the forceps jaws. Such a design permitsmaximum tissue engagement with small mechanical movement. As a result ofthe restricted access, and contrary to the teachings of the prior art,it is desirable to construct forceps with maximum treatment areas for agiven thermal mass or size.

[0069] One of the commonest design principles employed in bipolarendoscopic instruments is based on the Kleppinger forceps. Rather thanusing mechanical hinges, the opening of this type of forceps is achievedentirely by a spring force acting on the forceps jaws. Closure iseffected by sliding an outer tubular structure over the proximal springportion of the jaws. The forceps jaws are necessarily quite thin, so asto limit the forces needed to operate them. As a result, the jawsprovide negligible heat sinking for the given contact surfaces. Themechanical and biocompatibility properties of such tissue contact partsalso tend to result in the use of materials such as stainless steel,further reducing the capacity of the jaws to sink heat developed duringdelivery of bipolar r.f. energy. The jaws and the proximal sprung areacarry r.f. power, and the proximal portion is normally insulated using aplastics coating which further reduces the heat sinking capabilities.

[0070] Another exemplary forceps design based on the Kleppingeroperating principle is described in U.S. Pat. No. 5,445,638 (Rydell etal) and the commercial product based on this patent is sold by EverestMedical Corp., Minneapolis, USA as the BiCOAG Cutting Forceps. Thisforceps design includes the additional feature of a blade which may beadvanced along a space provided around the longitudinal axis of theforceps jaws such that, once the tissue pedicle is coagulated, it maythen be divided without needing a second instrument. The spacerequirement for operation of the blade yet further reduces the thermalmass and heat sinking capabilities of the forceps jaws. The opposingsurfaces of the forceps jaws commonly have teeth to prevent tissueslipping within the grasp of the instrument, particularly during theadvancement of the blade. For these teeth to provide simultaneouselectrosurgical and grasping functions, they have to mesh as they wouldbetween two gear wheels. This arrangement prevents the teeth frompiercing the tissue and shorting out the r.f. delivery. Unfortunately,the teeth have the effect of increasing the treatment area of thesurfaces of the forceps jaws, and increasing the thermal transfer fromthe tissue to the jaws. The best grasping function is achieved when theteeth are sharp, a feature that the prior art teaches against, as itincreases current density at the points of the teeth.

[0071] We aim to overcome these limitations in forceps or other bipolarelectrode designs by using high power pulses of bipolar r.f. energy toconvert the NTCR behaviour of tissue to a PTCR behaviour. A natural PTCReffect is realised by exploiting the current hogging phenomenon to theextreme.

[0072] Referring to FIGS. 4A to 4D, if high power is delivered, e.g. ata frequency in the range of from 100 kHz to 500 kHz, to tissue 10contained between the two contact surfaces (electrodes) 11 and 12 of abipolar instrument, current hogging ensues, as is illustrated in FIGS.4A and 4B. Thus, FIG. 4A illustrates initial power delivery to theelectrodes 11 and 12 with a low resistance region in the tissueresulting in uneven current density, and FIG. 4B illustrates theincreased current density which results from current hogging. If thepower is sufficiently high, then a vapour bubble 13 is formed within thetissue due to local temperature, as illustrated in FIG. 4C. This vapourbubble 13 contains pure steam which, at voltages substantially below300V peak, is completely insulating. The high current density across aregion of the tissue created by the current hogging phenomenon is,thereby, defeated by the insulative barrier of the vapour. The growth ofthe vapour bubble 13 is sustained by regions of high current density atregions 14 which occur at the periphery of the vapour bubble and arealong a line perpendicular to the current flow, as illustrated in FIG.4D. In effect the region of high current density is forced outwards byvapour propagation. If this growth in the vapour bubble 13 was allowedto continue, it would lead to an explosive popping, which could damagetissue outside the immediate application site. In fact, then, one of theprincipal factors limiting the power which can be applied using abipolar r.f. instrument/generator combination is steam bubble popping,an undesirable effect since it can prevent sealing.

[0073] By delivering high power only intermittently, sufficient time isallowed between activations to allow the vapour bubble to condensethereby to alleviate the pressure build up due to boiling ofelectrolytes. Another advantage of intermittent power delivery is thatthe clinical effect is slowed, ameliorating the difficulty in detectingand controlling the application of electrosurgical power to an optimumlevel. (For these reasons, power delivery in the prior art is usuallyrestricted to a rate consistent with an application time in the regionof five to ten seconds, with the result that prolonged application ofpower creates thermal damage adjacent the treatment site.)

[0074] An advantage of this technique is that current hogging to theextent of drawing significant current (due to a singular current hoggingpoint) is avoided. The preferred system produces multiple hot spotswithin a single burst, requiring the bipolar r.f. energy to be of a highcurrent which, typically for a 5 mm laparoscopic BiCOAG Cutting Forceps,has been found to be in excess of 1.5A; and, for a 10 mm version, up to4A.

[0075] Another benefit of high power bursts is that the thermalconduction from heated tissue to the forceps 11, 12 is limited. When thevapour bubble 13 is formed, there is a higher power density within thetissue than at the forceps/tissue interface. This higher power densityis the result of more protracted current pathways caused by multiplevapour bubbles. Tissue sub-surface to the forceps jaws, therefore, has ahigher effective resistivity. More power is delivered to sub-surfacetissue by virtue of higher voltage with less current, and so the tissueadjacent to the electrodes 11, 12 undergoes less heating. Duringexperimentation, tissue pedicles treated in this way show evidence ofdesiccation inside, but not on the surface. This finding is verydifferent from conventional bipolar r.f. electrosurgical power delivery,as the highest current densities normally occur at the tissue surface incontact with the forceps jaws.

[0076] The duty cycle of energy delivery can be adjusted to achieve thebest clinical effect. When energy is delivered to tissue in this way,the burst is of sufficient magnitude to cause vapour formation atmultiple sites within the tissue. In practical experimentation, thetissue is seen to swell with each burst as evidence of this. Powerdelivery then ceases before the vapour assumes a sufficiently highpressure to burst the tissue. The subsequent “off” period has to be longenough to ensure thermal relaxation. During this relaxation period,vapour recondenses, and aids the thermal conduction mechanism bycondensing preferentially at the coolest point. Moisture within thetissue is thus redistributed by this mechanism. The “off” time, theresultant thermal relaxation and the redistribution of moisture resultsin new current hogging points being created with each successive burst,ensuring an even distribution of effect in the tissue contained betweenthe electrodes 11, 12.

[0077] One of the difficulties associated with power delivery is therange of impedances encountered during use. Typical impedances can rangeanywhere between 10 ohms and 200 ohms. The maximum applied voltage islimited to a predetermined peak level which prevents arc propagationwithin the vapour. The peak voltage is, therefore, maintained below200V, e.g. using a voltage clamp circuit. For maximum power deliverywith this ceiling voltage, the waveform needs to be of low crest factor,typically less than 1.5. The most practical low crest factor waveform isa sine wave with a crest factor of 1.4. The maximum r.m.s. voltage is,therefore, 140V r.m.s. The maximum initial power delivery could,therefore, range between 100W and 2000W.

[0078] Instrument design can, however, limit maximum power delivery.Heating of the instrument as a result of resistive losses should to beavoided as far as possible. Generally, the thinner the tissue graspedbetween the forceps jaws of a given area, the lower the impedance. Thus,if the r.f. source behaved as a constant voltage source, power deliverywould be inversely proportional to tissue thickness. However, thinnertissue requires less energy to coagulate than thick. For example, if thetissue is half as thick, half the energy is required, yet power deliverywill be doubled. With a. constant r.f. voltage supply it is, therefore,desirable, to vary the duty cycle to reduce variation in the speed ofclinical effect, the speed of effect being proportional to the square ofthickness. It is possible that a particular instrument may be used overa 5 to 1 range of tissue thickness. Speed of effect variation would be25 to 1. The strategy of constant voltage and variable duty cycle isnot, therefore, preferred. The need to overcome current hogging in thintissue is greater than in thick tissue for the reasons outlined earlier.It has been found that a peak power of 200W is more than sufficient toachieve sub-surface vapour with the largest of instruments and thethinnest tissue. Limiting the power requirement rather than burstduration is advantageous in terms of instrument compatibility, reducingthe variation in treatment time and placing less demand on the r.f.generator. Changing burst duration whilst maintaining a constant r.f.voltage yields different treatment rates for different thicknesses.

[0079] The worst case for inducing sticking is when the tissue is thindue to the lack of current sharing, and this is often compounded by therequirements of the instrument design. As far as a single r.f. burst isconcerned, sufficient energy is supplied to create multiple vapourpockets. The energy requirement of the burst is determined by the volumeof tissue grasped, and hence the dimensions of the forceps jaws. Over awide range of instrument configurations, the energy requirement to reach100° C. may lie in the range of 2 to 20J. Minimum burst width at 200Wis, therefore, between 10 ms and 100 ms respectively. The latent heat ofvaporisation defines a corresponding energy requirement of 20 to 200J.This suggests that, if the burst is set to 200W for 100 ms, there wouldbe sufficient energy to vaporise the total electrolytes of the minimumtissue volume grasped. In practice, the sub-surface creation of vapourcauses a dramatic increase in impedance. The vapour formation and theabove-mentioned voltage clamping create an automatic regulating effectso that energy delivery beyond that needed to cause hot spots islimited. As the energy required for complete vaporisation is ten timesgreater, there is a large operating window of available settings. It is,therefore, possible to operate with a potential 20J of energy per burst.However, it is not necessary for this first burst to create vapour whenthe tissue is thick. The creation of vapour within thicker tissue has ahigher potential risk of popping. The auto-regulation of the maximumvoltage clamp reduces the burst energy into the higher impedancescreated by thicker tissue. Lower burst energy can, therefore, be usedthan that indicated in the earlier analysis, and yet still achieve thetissue effect. The auto-regulating effect is a function of the powerdelivery. The lower the burst power for a given energy, the lesspronounced this effect.

[0080] The subsequent “off”-time allows condensation and thermalrelaxation. This is a comparatively slow process. The hot vapourcondenses relatively quickly, but the subsequent thermal conduction isslow. Using forceps of low thermal mass and thermal conductivity, it hasbeen found that periods in excess of 100 ms are required beforesufficient thermal relaxation can take place. Values in the range of 300ms to 1s are preferred. This thermal relaxation is important to ensurethat the subsequent r.f. burst creates hot spots in previously untreatedareas of the tissue. The “on” time of each burst is typically in theregion of from 100 to 500 ms. These figures apply to power, voltage orcurrent waveforms, as do the mark-to-space ratio and duty cycle figuresreferred to in this specification.

[0081] The cycle of burst and relaxation times is continued until thetissue contained within the grasp of the forceps is completely treated.Due to the higher thermal capacity of thicker tissue, vapour may not begenerated in the first burst, but only in subsequent bursts. Electricalevidence of vapour generation is provided in the current and voltagetraces monitored during each power burst. When vapour is created, thevoltage clamp is reached and current decays. The next burst produces ahigher initial current as a result of the condensation during the“off”-time. This initial current is usually 50% greater than the endcurrent of the preceding burst. The auto-regulating effect of vapourcreation, in conjunction with voltage clamping, prevents completedesiccation. The current during each burst exhibits a decay similar toan exponential decay with the average value for each burst decreasing ina similar fashion. Vessel sealing occurs when the average deliveredcurrent decays to approximately 30% or less of its peak value. The mostnotable feature about the completion point is that the outer surfaces ofthe treated tissue in contact with the electrodes 11, 12 are stillmoist. The fact that this moisture is not vaporised helps prevent theextension of thermal damage beyond the treatment site which wouldotherwise occur as a result of surface steam condensing on adjacenttissue. The moisture also prevents tissue sticking, and the uniformityof treatment enables a more reliable determination of a coagulationend-point without the necessity of surface desiccation inherent inconventional systems.

[0082] FIGS. 5 to 11 illustrate the use of BiCOAG Cutting Forceps Foperated conventionally and as part of a system in accordance with theinvention. Each of these figures shows two perspective views of theforceps F, respectively from the distal end thereof and from the side.Thus, FIG. 5 shows a tissue pedicle P grasped in the jaws 21, 22 of theforceps F, the forceps being operated in a conventional manner. Thecurrent density between the forceps jaws (electrodes) 21, 22 is variableover the tissue contact area creating zones of high current density,shown by the arrowhead symbols 23 in FIG. 5. The variations inimpedances which may occur as a result of, amongst other things, thenon-parallel closure of the forceps F creates the zones 23 of highcurrent density. The zones 23 of high current density create hot spotsat the contact surfaces between the tissue and the forceps jaws 21, 22.The hot spots created in the zones 23 of high current density reduce theimpedance of these zones even further compared to the other areas of thetissue. All the current from the output becomes concentrated in thesehot spots which exhibit the phenomenon of current hogging. The hot spotsbecome even hotter until the tissue on the surface becomes completelydesiccated and the impedance falls. Only then will the areas ofuntreated tissue then be treated. This is well demonstrated when theproximal end of the forceps jaws are more closely opposed than the tips,in that the effect is seen to move along the length of the jaws duringapplication. Current hogging produces two undesirable effects: thetissue surface must be desiccated to ensure complete treatment whichincreases the risk of tissue sticking, and the application time must beprolonged to ensure complete treatment which increases the collateralmargin.

[0083] The generator and system described in this specification overcomethese problems in the following ways. The zones of high current densityare instantly created, as shown in FIG. 1, by the burst of bipolar r.f.energy. As has already been described, these zones of higher currentdensity are more likely to be created in thinner tissue when the forcepsjaws are more closer together. This situation can be created by firstgrasping the tissue within the jaws, and preferably employ a ratchetfeature on the BiCOAG Cutting Forceps so that the tissue is crushed andheld at an optimal cross-section. Under these circumstances, when thefirst pulse is applied, the tissue in the zones of high current densityreaches 100° C. virtually instantly.

[0084] FIGS. 6 to 11 show the use of the forceps F when operated as partof a system in accordance with the invention, that is to say the forcepsF are supplied with electrosurgical energy by an r.f. generator asdescribed in this specification. Thus, as shown in FIG. 6, the power ofthe first pulse is dissipated in the centre of the tissue pedicle P inzones 23 of high current density, creating pockets 24 of water vapour(steam) in the intracellular and interstitial fluids. High current andhigh power are used to form the vapour pockets 24.

[0085] Such power and current levels are not normally available from aconventional bipolar electrosurgical generator for “dry field”electrosurgery. The creation of the vapour pockets 24 produce twobenefits: the vapour pockets 24 produce a high impedance barrier whichprevents further current hogging, and the highest current densitiesoccur around the lateral edges of the vapour pockets, as shown in FIG.7. Heat generation and coagulation start internally, within the tissuepedicle P, rather than in the external contact area between the tissueand the forceps jaws 21, 22.

[0086] Referring now to FIG. 8, the pathway of least resistance for thecurrent flow is around the vapour pockets 24. This concentration ofcurrent expands the vapour pockets 24 at their lateral edges where thehighest temperatures occur. The tissue effect, therefore, naturallymoves to untreated areas within the pedicle P. During use, the tissue isseen to swell with each energy pulse. If, however, the vapour were topersist in growing, less and less tissue would be conducting thecurrent. This would generate vapour far more quickly, so that apotential runaway situation could occur, producing the bursting orpopping associated with prolonged application from a conventionalgenerator. The auto-regulating feature of the present system shuts offthe power of a given energy pulse in microseconds when excessive vapourformation occurs. Excessive vapour formation is further avoided by thetermination of the energy pulses in accordance with the cycle of burstand relaxation times mentioned above.

[0087] Referring now to FIG. 9, when the first energy pulse isterminated, the vapour pockets 24 collapse, leaving areas 25 ofdesiccation inside the tissue pedicle P but none on the surfaces betweenthe forceps jaws 21, 22 and the pedicle, which surfaces remain moist.Heat generated within the tissue pedicle P dissipates in the colderareas of the pedicle as the vapour condenses. Once this thermalrelaxation has been allowed to occur, a second energy pulse is appliedas shown in FIG. 10. The zones 23 of high current density are nowcreated in previously untreated areas, because of the higher impedanceof the desiccated tissue produced by the first energy pulse. Vapourpockets 24 (not shown in FIG. 10) once again form in these zones, andexpand laterally to include any untreated areas.

[0088] The on-off cycle of bipolar r.f. energy pulses is continued untilthe power absorption at each pulse falls below a level indicative ofcomplete coagulation, as indicated by the reference numeral 26 in FIG.11. This point corresponds to the point at which no more zones of highcurrent density can be created. This gives an automatic indication whenthe tissue within the pedicle P is uniformly treated with surfacecoagulation, but not desiccated. The maximum effect is produced withinthe tissue pedicle P with the surfaces adjacent to the jaws 21, 22remaining moist and non-adherent to the jaws.

[0089] The measures described above provide for faster uniformcoagulation of vascular pedicles without the need to skeletonise.Skeletonisation is a surgical technique in which the fat and connectivetissue which normally surrounds vessels is removed to expose the vesselsthemselves. This removes what, in effect, is a high impedance barrier tothe transfer of bipolar r.f. energy to the lower impedance vascularstructures within a pedicle. The advantage of the present system in thissituation is provided by the preferential absorption of energy within apedicle.

[0090] During practical use of the system, a surgeon will need todeliver less energy to achieve a therapeutic effect than if aconventional, continuous, bipolar r.f. output was used. The graph ofFIG. 12 illustrates the therapeutic effect on tissue after delivery of acertain amount of energy over a certain amount of time delivered from acontinuous output bipolar r.f. source. During an initial phase 27 of atreatment cycle, energy delivery is effective. As current hoggingoccurs, some tissue areas reach the therapeutic level before others. Tocreate haemostasis, all tissue areas need to achieve this level. Toensure that these other regions are brought to the therapeutictemperature, power has to be applied for a longer period of time. Duringthis extension period 28 of the treatment cycle, most of the appliedenergy is wasted in boiling the electrolytes in the region thatinitially formed the current hogging point. The appropriate treatmenttime is often so indeterminate that power is applied until completedesiccation occurs. Boiling occurs while power is maintained at a presetlevel 29. Once desiccation occurs, the load impedance rises and thedelivered power decreases, as shown by the decay part 30 of the curve inFIG. 12. This excessive boiling of electrolytes helps explain tissuesticking, charring and lateral thermal margins.

[0091] The graph of FIG. 13 illustrates how a desired therapeutic effectcan be reached using the present system after three r.f. pulses 31, 32and 33 are applied. Each pulse 31, 32, 33 is followed by a respective‘relaxation period’ 31 a, 32 a, 33 a. The first pulse 31 that is appliedis capable of creating vapour. As this vapour forms internally, itinterferes with power delivery, causing a reduction in power (indicatedby the line 31b) towards the end of the pulse. The energy absorbed byvaporising the small quantity of electrolyte involved is thenredistributed during the ‘relaxation period’ 31 a before the next pulse32 of energy is delivered. This redistribution occurs by condensation.The amount of vapour produced by each subsequent pulse is greater, andso results in even further power reductions, but also an even greaterdispersion of energy throughout the tissue. This redistribution ofenergy by the condensing vapour is demonstrated by the fact that theinitial energy delivery for each pulse is not interrupted by vapour. Theenergy of each pulse, as represented by the shading in FIG. 13 is almostentirely effective. As little or no excess energy is used, and theheating occurs from inside to outside (unless the surgeon choosesotherwise, e.g. when a thermal treatment margin is required), there willbe little excess thermal energy to cause sticking, charring orcollateral tissue damage. The graphs of FIGS. 12 and 13 can be obtainedby application of a pair of forceps to morbid vascular tissue andenergising continuously or in pulses respectively.

[0092] Referring to FIG. 14, an electrosurgical system in accordancewith the invention comprises a generator 40 for generating radiofrequency power, and an electrosurgical instrument comprising theassembly of a handheld forceps unit 42, a connecting cable 44, and aconnector 46 for removably connecting the assembly to the generator 40via a generator connector 48 containing the generator output terminals.Instead of being on the generator, the connection interface between theforceps unit 42 and the generator 40 may be on the forceps unit 42itself, the significant point being that alternative treatment units,whether forceps or otherwise, may be connected to the generator 40.

[0093] The forceps unit 42 has a pair of electrodes 50 which are coupledvia power delivery conductors 52 passing through the body of the forcepsunit 42 and the cable 44 to the connector 46 where they are connected totwo of the output terminals (not specifically shown) of the generator ingenerator connector 48, to allow supply of radio frequency power fromthe generator to the electrodes. Radio frequency power for supply to theelectrodes 50 is generated in an r.f. output stage 60 having outputlines 62 associated with respective output terminals in the generatorconnector 48. As described above, the generator 40 is arranged to supply100% amplitude-modulated radio frequency power with a carrier frequencyin the range of from 100 kHz to 500 kHz and with a pulse repetition ratein the region of 0.7 to 3 Hz, typically. The modulating waveform is fedto the r.f. output stage 60 by a pulse modulator 64 via connection 66.

[0094] The peak r.f. voltage generated between the output stage outputlines 62 is limited, typically to 200V peak, by the combination of avoltage threshold detector 68, coupled between the lines 62, and acontroller stage 70. When the voltage threshold, set by the controllervia threshold set line 72, detects a peak output voltage exceeding theset threshold voltage, a threshold detect signal is fed to thecontroller 70 via the detector output line 73 and the r.f. power isreduced by adjusting a switched mode power supply 74 which suppliespower to the output stage 60, the controller signal being applied viapower set line 76.

[0095] Another function of the controller 70 is to set the frequency andpulse width or mark-to-space ratio of the pulse modulation applied tothe r.f. output stage 60 by the pulse modulator 64.

[0096] The controller 70 also receives an output current detectionsignal from a current detector circuit 77 coupled in one of the outputlines 62 by a current transformer 78.

[0097] It will be appreciated that when, during use of the system, thesurgeon wishes to coagulate, for instance, a pedicle, between theelectrodes 50 of the forceps unit 42, he operates the forceps to graspthe pedicle between the electrodes 50 and activates the generator 40 bymeans of a foot switch (not shown), whereupon the r.f. output stage 60is activated by the pulse modulator 64 so that a 100%amplitude-modulated r.f. signal is fed to the electrodes 50 at afrequency set by the controller 70, the mark-to-space ratio being suchthat the “off”-time of the output stage 60, as determined by thecontroller 70 and the pulse modulator 64, is at least 100 ms betweeneach successive pulse. With successive pulses, the instantaneous appliedpower decays towards the end of each pulse as vapour is formed withinthe tissue. As described above, the “off”-times 31a, 32a, 33a (FIG. 13)are each sufficient to allow the vapour within the tissue to condensebefore application of the next pulse, but in each successive pulse, ifthe applied power level is held constant, the power decays to afinishing value lower than that occurring in the previous pulse. In thepresent embodiment, this decay is sensed by the current detector circuit77 and the controller 70, and the controller is arranged to terminatethe pulses when the rms current at the end of one of the pulses fallsbelow a predetermined fraction of the rms current at the beginning ofthe pulse. In this case, the pulses are cut off when the finishingcurrent is 30% or less than the starting current. Accordingly, in thisembodiment, a current threshold is used to terminate a sequence ofpulses, i.e. termination occurs when the r.f. current falls below apredetermined current threshold. As an alternative, the sensingcircuitry of the generator 40 may be arranged to deliver a sensingsignal to the controller which is proportional to power, so thattreatment can be terminated when the instantaneous power falls below apredetermined power threshold. Variations on this principle may be used,including current or power thresholds which are absolute, or which arespecified as a fraction of a value at the commencement of treatment, oras a fraction of the value at the commencement of the pulse in question.

[0098] At this point it is worth noting that the combination of thepulsed output and a voltage limit (typically 120V rms) create a powerversus impedance load curve (averaged over the pulses) which is somewhatnarrower than that of conventional generator operating with a continuousoutput. This is illustrated in FIG. 15. The present generator maytypically produce an instantaneous power output of 200W when running ata 15% duty cycle, the current being limited to a value in the region of1 amp to 5 amps rms, which yields a power peak between 10 ohms and 100ohms load impedance, in contrast to a conventional generator operatingat an average power of, typically, 30 watts which would produce anapproximately flat power-versus-load impedance curve in which power ismaintained at or near a maximum value over a ten-fold range ofimpedance, e.g. from 10 ohms to well in excess of 100 ohms. In FIG. 15,the dotted curve A corresponds to a 15% duty cycle pulsed output with apeak power output of 200W and a current rating of 4 amps r.m.s. Thesolid curve B represents the power-versus-impedance characteristic forthe conventional generator operating with continuous r.f. output of 30W.Both curves are voltage-limited at 120V r.m.s. It will be seen thatalthough the pulsed generator delivers its maximum power over a narrowerimpedance range than the continuously operating generator, neverthelessmaximum power is delivered over a load impedance range starting at nomore than 20 ohms. A realistic lower limit for peak power delivery is100W when driving loads down to 20 ohms, recognising that the maximumimpedance into which this peak power can be delivered can be deliveredis determined by the voltage limit (here 120V r.m.s) imposed to preventarcing. The limitation in load curve width is desirable inasmuch as itprovides the auto regulation feature described above at the end of thetreatment. The extent to which power can be delivered into a low loadimpedance is governed by the current rating of the generator. In thepresent generator, an rms current value in excess of 1.5 amps at thestart of each pulse is typically achieved, with 3 or 4 amps beingattainable.

[0099] It will be appreciated that if the electrodes 50 of the forcepsunit 42 are comparatively large in their tissue contact area, the loadimpedance presented to the generator will be comparatively low. The loadimpedance also decreases as the thickness of tissue grasped between theelectrodes 50 decreases. It is possible to improve the speed oftreatment by altering the pulses produced by the generator according tothe characteristics of the instrument to which it is connected. Althoughlarge area electrode produce a low load impedance, the thermalrelaxation time of the larger area of tissue grasped is longer due tothe longer thermal conduction paths. Smaller area electrodes can betreated with a larger duty cycle or mark-to-space ratio, due to thelower thermal relaxation times, and with lower peak power. Larger dutycycles have the effect of increasing the ability of the generator tomatch into high impedance loads (due to the power-versus-load peakextending to higher impedance values). Consequently, increasing the dutycycle when the electrodes are small in area provides the advantage offaster treatment.

[0100] Changing the pulse duty cycle, then, in conjunction with theupper voltage clamp has the effect of changing the load curve to suitthe instrument being used. Referring again to FIG. 14, pre-adjustment ofthe pulse characteristics may be performed by arranging for theinstruments which are to be connected to the generator 40, such asforceps unit 42, to have an identification element 80 which may besensed by a sensing circuit 82 in the generator when the instrument isconnected to the generator output connector 48. In the example shown inFIG. 14, the identification element 80 is a capacitor of a specificvalue coupled between one of the power leads 52 and a third lead 84 inthe cable. These same two leads are coupled via the connectors 46, 48 toa pair of inputs 86 of the sensing circuit 82, which acts as anelectrode identifying circuit by responding to the value of thecapacitor. The controller 70 varies the basic or initial pulse dutycycle according to an identification signal received from theidentification circuit 82 via line 88. Details of the electrodeidentification circuit 82 and its interaction with the identifyingelement 80 are described in European Patent Publication 0869742A, thecontents of which are incorporated herein by reference.

[0101] Accordingly, by arranging for different value capacitors 80 to beincorporated in different instruments according to, for instance,electrode tissue contact area and other properties of the instrumentaffecting load impedance and thermal relaxation time, the generator canbe automatically configured to produce a pulsed output particularlysuited to the instrument in question. In particular, as instruments withlarger tissue contact areas are selected, the preset duty cycle ormark-to-space ratio is lowered and/or the pulse frequency is lowered.

[0102] The controller may alter not only the mark-to-space ratio, butalso pulse frequency and power output via, in this case, the pulsemodulator 64 and/or the switched mode power supply 76.

[0103] As an alternative to identifying the instrument or instrumentcategory, the generator 40 may be provided with a sensing circuit forsensing the load impedance across the output lines 62 of the outputstage 60 at or around the instant at which the surgeon commenceselectrosurgical treatment, the pulse characteristics thereby set beingmaintained until treatment is finished. Referring to FIG. 15, the pulseduty cycle can be increased, as shown, with increased initial loadimpedance. In the example shown, the duty cycle is maintained below 50%(i.e. a mark-to-space ratio of 1:1) for impedances less than about 140ohms. Referring to FIGS. 15 and 16 together, the controller may bearranged, in addition, to set the peak power (FIG. 15) and the pulsefrequency (FIG. 16) concurrently according to the initial loadimpedance, the power being set higher and the pulse frequency being setlower for low initial impedances than for high initial impedances.

[0104] The initial impedance may be sensed by monitoring the current,given that for a known initially applied power, the initial loadimpedance is inversely proportional to the square of the output current.

[0105] Further benefits can be obtained by arranging the generator so asto perform a treatment cycle consisting not merely of a plurality ofpulses of a single preset duty cycle, but by dividing the treatmentcycle into periods in which the generator output signal begins as apulsed r.f. signal with a predetermined duty cycle and finishes with adifferent characteristic. Referring to the power-versus-time graph ofFIG. 18, the treatment cycle may have an initial period 130 in which ther.f. power signal consists of a series of pulses 131, 132, 133 with apredetermined duty cycle, followed directly by a subsequent period 140in which the r.f. power signal is a c.w. signal 141 of much lower poweramplitude. Typically, during the initial period 130, the pulses 131 to133 have a duty cycle in the region of from 15% to 30% with a peak powerof 200W. The transition from the initial period 130 to the subsequentperiod 140 may be controlled by feedback from the output circuitry ofthe generator. Referring back to FIG. 14, the switched mode power supply74 is controlled via line 76 by the controller 70 which is, in turn,responsive to a sensing signal on line 73 from the output voltagethreshold detector 68. Being a switched mode device, the power supply 74has its own switching frequency which, in this embodiment, may be in theregion of 25 kHz, supplied as a pulse stream from the controller 70. Inthis example, the r.f. output voltage of the generator 40 is limited byinterrupting the switching pulses supplied to the power supply 74 whenthe output voltage exceeds a predetermined threshold (typically 120Vr.m.s., as mentioned above). By monitoring the power supply switchingpulses generated by the controller 70, it is possible to determine theamount of energy delivered by the generator. Counting the switchingpulses, therefore, offers a convenient way of monitoring electricalconditions at the generator output. In particular, referring to FIG. 18,the decrease in delivered power due to the formation of vapour in thetissue and visible as decay curves 132 b and 133 b in the powerwaveform, is the result of interruptions in the power supply switchingpulses produced in response to the output voltage having exceeded thethreshold set in the voltage threshold detector 68 (FIG. 14).Accordingly, by counting the power supply switching pulses, it ispossible to determine when the low duty cycle waveform ceases to beadvantageous, whereupon the controller 70 can adjust its output to causethe switched mode power supply to deliver energy on a continuous or morenearly continuous basis, but at a significantly lower peak power level,as illustrated by the c.w. waveform 141 in FIG. 18. Typically, theaverage power delivered during this subsequent period 140 of thetreatment cycle is the same as the average power delivered during theinitial period 130.

[0106] The more nearly continuous power delivery may be obtained,instead, by arranging for the r.f. power signal during the subsequentperiod to take the form of a pulsed signal with a significantly higherduty cycle but lower peak power, as shown in FIG. 19. In thisembodiment, the controller 70 is arranged such that, as before, theinitial period 130 of the treatment cycle consists of a plurality ofpulses 131, 132, 133 with a low duty cycle and high peak power. Again,the transition to the subsequent period 140 is carried in response toelectrical conditions at the generator output. In the subsequent period140, however, the duty cycle is higher, e.g. at least twice that of theinitial period, and the peak power is correspondingly reduced to resultin at least approximately the same average power. Further vapourformation may occur in the tissue during the subsequent treatment cycleperiod 140 resulting in operation of the voltage clamp in the same wayas during the initial period, as evident from the decay portion 143 b ofpulse 143.

[0107] In an alternative embodiment, not shown in the drawings, thetreatment cycle may have more than two periods in which the r.f. powersignal has different characteristics. In particular, the signal mayconsist of a succession of pulses beginning with a first group of pulseshaving a first low duty cycle, followed by a second group of pulseshaving a second greater duty cycle, followed by a third group of pulseswith a third, yet greater duty cycle, and so on, so as to maintainoptimum coagulation effectiveness as the tissue characteristics change.In other words, a three stage treatment cycle may be employed, eachstage consisting of a number of pulses with its own respective fixedduty cycle. Typically, the successive stages have pulses with dutycycles of 15%, 30%, and 60%, and peak power values of 200W, 100W and 50Wrespectively, in order to maintain an approximately constant averagepower delivery.

[0108] The effect common to all three alternatives described above isthat the load curve of the generator has an initially narrowcharacteristic as exemplified by curve A in FIG. 15, but is extended inthe high impedance range as energy delivery becomes more nearlycontinuous, whether in the form of a c.w. output 141 as in FIG. 18, orin the form of an output with a higher duty cycle, as in FIG. 19. Itfollows that coagulation of the tissue being treated proceeds morequickly since power delivery into the tissue is maintained as the tissueimpedance increases owing to vapour formation and, subsequently,localised coagulation of tissue.

[0109] Referring back to FIG. 13, it will be recalled that the“off”-times 31 a, 32 a, 33 a are each sufficient to allow the vapourwithin the tissue to condense before application of the next pulse, butin each successive pulse, if the applied power level is held constantthe power decays to a finishing value lower than that occurring in theprevious pulse. Referring again to FIG. 14, this decay can be sensed bythe current detector circuit 77 and the controller 70, and in a furtherembodiment of the invention the controller is arranged to reduce theapplied power amplitude and increase the width of any given pulse in thesuccession of pulses when the rms current at the end of the precedingpulse falls below a predetermined fraction of the rms current at thebeginning of that preceding pulse. For instance, the pulse width may beincreased and the power amplitude decreased when the finishing currentis 30% or less than the starting current. Accordingly, in thisembodiment, a current threshold is used to control a progressiveincrease in pulse width and mark-to-space ratio and a progressivedecrease in pulse amplitude.

[0110] In this preferred embodiment, the pulse power amplitude of thesuccession of pulses is progressively decreased and the pulse widthprogressively increased according to the time taken from the pulserising edge to the instant at which the voltage threshold is reached, itbeing understood that the instantaneous power decay at the end of eachpulse shown in FIG. 13 represents a constant voltage portion of thepulse, the voltage threshold imposed by the threshold detector 68 ofFIG. 14 having been reached. Accordingly, by setting a time thresholdfor the initial portion of the pulse, i.e. the time delay to reach thevoltage threshold, it is possible to cause the next pulse to be reducedin power amplitude and increased in width when the time delay is lessthan the time threshold. Variations on this principle are possible. Forexample, the degree to which the time delay is less than the timethreshold may be used to control the degree to which the next pulse isreduced in amplitude and increased in width. Alternatively, the fractionof the respective pulse width which elapses before the voltage thresholdis reached may be used to control the succeeding pulse amplitude andwidth, rather than having a fixed time threshold.

[0111] In another preferred embodiment, the sensing circuitry of thegenerator 40 may be arranged to deliver a sensing signal to thecontroller which is proportional to load resistance, so that pulseamplitude can be decreased and pulse width can be increased as the loadresistance rises. Variations on this principle may be used, includingcurrent or load impedance thresholds which are absolute, or which arespecified as a fraction of a value at the commencement of treatment, oras a fraction of the value at the commencement of the pulse in question.

[0112] The tailoring of the pulse width or duty cycle to smaller orlarger area electrodes can be achieved at least to some extentautomatically by controlling the change in pulse amplitude and widthaccording to load resistance as set out above, but additional matchingof pulse characteristics to instruments differing in this way can beobtained by setting the initial amplitude and width according to anidentification element forming part of the instrument and sensed by thegenerator when the instrument is connected.

[0113] In particular, as instruments with larger tissue contact areasare selected, the starting duty cycle or mark-to-space ratio can belowered and/or the pulse frequency lowered.

[0114] The effect of controlling pulse amplitude and pulse width in theways described above is illustrated by way of example by the modulationwaveform shown in FIG. 20. The generator output signal begins as apulsed r.f. signal with a predetermined initial pulse width and dutycycle and finishes with a different characteristic, the pulses changingprogressively in between. Thus, the treatment cycle may have an initialpulse 150 the width of which is between 15% and 30% of the pulse period.This is followed by pulses 151, 152, 153, 154 of progressivelydecreasing amplitude and increasing width as the impedance of the loadincreases. The initial pulse 150 typically has a peak power of 200W.

[0115] The progressive change in amplitude and width may be controlledby feedback from the output circuitry of the generator. Referring backto FIG. 14, the switched mode power supply 74 is controlled via line 76by the controller 70 which is, in turn, responsive to a sensing signalon line 73 from the output voltage threshold detector 68. Being aswitched mode device, the power supply 74 has its own switchingfrequency which, in this embodiment, may be in the region of 25 kHz,supplied as a pulse stream from the controller 70. In this example, ther.f. output voltage of the generator 40 is limited by interrupting theswitching pulses supplied to the power supply 74 when the output voltageexceeds a predetermined threshold (typically 120V r.m.s., as mentionedabove). By monitoring the power supply switching pulses generated by thecontroller 70, it is possible to determine the amount of energydelivered by the generator. Counting the switching pulses, therefore,offers a convenient way of monitoring electrical conditions at thegenerator output. In particular, referring to FIG. 20, the decrease indelivered power due to the formation of vapour in the tissue and visibleas decay curves 150 b, 151 b, and so on in the power waveform, is theresult of interruptions in the power supply switching pulses produced inresponse to the output voltage having exceeded the threshold set in thevoltage threshold detector 68 (FIG. 14). Accordingly, by counting thepower supply switching pulses, it is possible to determine the rate atwhich the pulse amplitude should be decreased and pulse width increased,these changes in output being effected by the controller 70. Typically,the average power delivered during each pulse remains substantiallyconstant over at least a major part of the treatment cycle from pulse150 to the cessation of power application.

[0116] As mentioned above, the rate at which the pulses increase inwidth and decrease in amplitude can be determined, instead, by sensingload resistance. All in all, by arranging for the pulses to changeprogressively in this way, optimum coagulation effectiveness may bemaintained as the tissue characteristics change. In other words, acontinuously varying treatment cycle may be employed, consisting of anumber of pulses each with its own respective width and amplitude.Typically, the pulse duty cycle increases from 15% to at least 60% andpeak power values of 200W, down to 50W or less are used to maintain anapproximately constant average power delivery.

[0117] Again, the effect of a modulation waveform such as that describedabove with reference to FIG. 20 is that the load curve of the generatorhas an initially narrow characteristic as exemplified by curve A in FIG.15, but is extended in the high impedance range as energy deliverybecomes more nearly continuous, whether finally in the form of a c.w.output, or in the form of an output with a higher duty cycle.

[0118] In another embodiment of the invention, the treatment cyclebegins with at least one pulse of a predetermined ON-time (i.e.duration) followed by a predetermined OFF-time (i.e. duration ofsubstantially zero power delivery). Thereafter, in a subsequent periodof the treatment cycle, the r.f. power signal has a differentcharacteristic, the transition from the initial period to the subsequentperiod being controlled by the control circuitry in response to the atleast one monitored parameter. Preferably, the ON-time is less than orequal to 300 ms and the OFF-time is longer than the ON-time. As typicalvalues, the ON-time is in the region of 100 ms and the OFF-time is about700 ms. This implies a pulse repetition rate when plural pulses aredelivered, in the region of 1 pulse per second. Alternative pulserepetition rates are possible, such that, for instance, the pulserepetition rate is 140 ms with the ON-time in the region of 40 ms andthe OFF-time in the region of 100 ms.

[0119] The subsequent period of the treatment cycle may be characterisedby a c.w. signal of a lower instantaneous power amplitude than theinstantaneous power amplitude of the pulse or pulses of the initialperiod. Alternatively, the subsequent period may be characterised by aplurality of pulses of fixed or variable mark-to-space ratio and/orfixed or variable amplitude, the amplitude being different (preferablyless than) the amplitude of the pulse or pulses of the initial period.

[0120] Preferably, the monitoring of an electrical parameter associatedwith the output terminals is performed during at least the first pulseoccurring during the initial period. Accordingly, in the case of themonitored parameter being the time from the beginning of the pulse tothe instant at which the output voltage reaches the above-mentionedvoltage threshold, the next pulse may have the same ON-time as the firstpulse, or the above-mentioned subsequent period may be commenced. Inother words, according to whether a predetermined condition associatedwith the monitored electrical parameter occurs in the first or asubsequent pulse of the initial period, the treatment cycle may consistof one pulse of predetermined duration or a plurality of pulses havingthe same predetermined duration, each pulse being followed by anOFF-time of a predetermined length.

[0121] Other electrical parameters may be monitored to determine whenthe initial period of the treatment cycle ends, for example, theabsolute or fractional difference between the instantaneous poweramplitude achieved prior to the output voltage reaching the presetvoltage threshold and the instantaneous power amplitude at the end ofthe pulse, or the sensed load resistance. In either case, the parameteris monitored during each pulse of the initial period in order todetermine when the pulsing pattern of the initial period (i.e.characterised by the predetermined ON-time and predetermined OFF-time)should be terminated.

[0122] One advantage of the invention is illustrated when attempting tocoagulate vessels immersed in blood or other conductive fluid. Withconventional bipolar generators, the presence of blood causes current tobe dissipated into the blood rather than into the tissue or bleedingvessel. This is due to blood conducting the current better than thetissue between the two jaws, a situation which will create currenthogging. This means that, to achieve haemostasis, the current must beapplied for a long period of time, thereby ensuring hot spots, charringand sticking. Using the present system, the bipolar r.f. energy pulsesare applied for very short periods of time, and the formation of vapourprevents the current hogging. This ensures that the tissue receivessufficient energy to achieve haemostasis, and is not preferentiallydissipated into the blood as a result of hot spots.

[0123] The features of the present system are particularly useful whenperforming endoscopic surgery, wherein vascular structures requiredivision or dissection in a bloodless fashion. Typical proceduresinclude laparoscopic procedures, such as laparoscopic assisted vaginalhysterectomy and laparoscopic supracervical hysterectomy wherein theuterine and other associated vessels require division; laparoscopicNissen fundoplication, where the short gastric and other associatedvessels require division; laparoscopic procedures on the bowel, whereoften the mesenteric vessels require division; laparoscopicappendicectomy, where the appendiceal artery and other associatedvessels require division; mobilisation of the omentum where the omentalvessels require division; laparoscopic bipolar tubal ligation, where thefallopian tube is coagulated to induce sterility; and, in general, forthe division of vascular adhesions. In all cases, the cauterisation canbe achieved without protracted dissection of the vascular structures toskeletonise them prior to sealing and division.

[0124] Other exemplary endoscopic procedures include minimal accesscardiac surgery, where vascular structures (such as the internal mammaryartery or gastroepiploic artery) are mobilised by division of branchesprior to bypass; and the harvesting of other vascular structures (suchas the saphenous vein) where once again the tributaries requiredivision.

[0125] The present invention is not restricted to use with bipolarforceps. It may be used to advantage in other bipolar instrumentation toeffect coagulation. The two poles of such an instrument, such as bipolardissecting hooks, are often in close proximity, such that any conductivematerial between the hooks creates the shortest conductive path withlimited penetration of the energy to the tissue against which theinstrument is applied. By interrupting the current path directly betweenthe hooks, as a result of vapour formation, a greater effect may beobtained in the tissue compared to conventional outputs.

[0126] Open surgical instruments such as bipolar forceps and the likemay be used.

[0127] In the preferred embodiment described above, the application ofr.f. power to produce the desired clinical effect is performed with aminimum burst energy capable of creating vapour within the graspedtissue. In particular, the burst energy is high enough to create vapourfrom the first burst when tissue is thin. This energy is delivered at apower sufficiently high that voltage clamping takes place within theburst, a thermal relaxation time before the next burst of at least 100ms being allowed.

What is claimed is:
 1. An electrosurgical generator comprising a sourceof radio frequency (r.f.) energy, at least a pair of output terminalsfor connection to a bipolar electrosurgical instrument and fordelivering r.f. energy from the source to the instrument, and a pulsingcircuit for the source, wherein the pulsing circuit and the source areso arranged as to deliver into a resistive load, when connected acrossthe output terminals, an amplitude-modulated r.f. power signal in theform of a succession of pulses, which succession is characterised bysuccessive pulses being of progressively decreasing power amplitude andof progressively increasing pulse width.
 2. A generator according toclaim 1, wherein the increase in the pulse width of successive pulses isrelated to the decrease in amplitude.
 3. A generator according to claim2, wherein the increase in pulse width and decrease in power amplitudeis such that the energy delivered in each pulse is substantially fixedfrom pulse to pulse.
 4. A generator according to claim 1, includingcontrol circuitry including means for monitoring at least one electricalparameter associated with the output terminals, the decrease in thepulse amplitude and the increase in the pulse width being effected inresponse to the at least one electrical parameter.
 5. A generatoraccording to claim 4, wherein the at least one parameter is theresistance of the load.
 6. A generator according to claim 4, wherein thegenerator includes circuitry for limiting the voltage applied to theload to a predetermined maximum value, and wherein the at least oneparameter is the time from the start of the pulse for the voltage toreach the said predetermined maximum value.
 7. A generator according toclaim 6, wherein the predetermined maximum voltage value is below 200Vpeak.
 8. A generator according to claim 1, wherein the pulse repetitionrate is less than 1 Hz.
 9. A generator according to claim 1,characterised in that the pulses are separated by periods of at least100 ms in which the delivered power is substantially zero.
 10. Agenerator according to claim 1, wherein the initial pulse width is inthe range of from 25 ms to 750 ms.
 11. A generator according to claim 1,wherein the succession of pulses is terminated by a period of continuousenergy delivery.
 12. A generator according to claim 11, wherein theperiod of continuous energy delivery is at least 1 second in duration.13. A generator according to claim 1, wherein the depth of amplitudemodulation in the said r.f. signal is at least 90% and the initial pulsemark-to-space ratio is less than 2:3.
 14. A generator according to claim1, arranged such that the pulsing frequency remains constant at apredetermined value over the major part of the said succession ofpulses.
 15. A generator according to claim 1, arranged such that atleast one of the initial pulse mark-to-space ratio, the initial pulsewidth, and the pulsing frequency is automatically preset at the start ofthe said succession of pulses to a value which is dependent on aninstrument identification signal received by the generator.
 16. Agenerator according to claim 1, arranged such that at least one of thepulse mark-to-space ratio, the pulse width, and the pulsing frequency isautomatically preset at the start of the said succession of pulses to avalue which is dependent on the load resistance cross the outputterminals at the said start.
 17. A generator according to claim 1,arranged so as to be capable of delivering a peak power of at least 100Winto any resistive load connected across the generator output in therange of from 20 ohms to 250 ohms.
 18. An electrosurgical generatorcomprising a source of radio frequency (r.f.) energy, at least a pair ofoutput terminals for connection to a bipolar electrosurgical instrumentand for delivering r.f. energy from the source to the instrument, and apulsing circuit for the source, wherein the pulsing circuit and thesource are arranged such that the r.f. energy is delivered over a periodof time constituting a treatment cycle in which, when a resistive loadis connected across the output terminals, is delivered as anamplitude-modulated r.f. power signal in the form of a succession ofpulses in which the periods between successive pulses are at least 100ms, the depth of amplitude modulation is at least 90% and the pulsemark-to-space ratio increases progressively from a value less than 2:3at the start of the treatment cycle.
 19. A generator according to claim18, arranged such that the peak power decreases progressively as themark-to-space ratio increases.
 20. A generator according to claim 19,arranged such that over at least a major part of the said succession ofpulses, the energy delivered per pulse remains substantially constant.21. A generator according to claim 18, wherein the pulse repetition rateis less than 1 Hz.
 22. A generator according to claim 21, wherein theperiods between successive pulses of the succession of pulses is 300 msor longer.
 23. A generator according to claim 18, arranged such that thepeak voltage of the said r.f. signal remains below 200V peak when aresistive load is connected across the output terminals.
 24. A generatoraccording to claim 18, wherein the pulsing circuit is arrangedautomatically to adjust the mark-to-space ratio of the signal pulses inresponse to a sensing circuit associated with the output terminals. 25.A generator according to claim 24, wherein the sensing circuit isarranged to be responsive to an identification element housed in aninstrument connected to the output terminals.
 26. A generator accordingto claim 22, in which the sensing circuit is arranged to detect aninitial value of a load impedance associated with the start of thetreatment cycle.
 27. A generator according to claim 26, wherein thepulsing circuit is arranged such that the initial pulse mark-to-spaceratio increases with increasing sensed initial load impedance.
 28. Amethod of electrosurgically coagulating tissue between the electrodes ofa bipolar electrosurgical instrument, in which r.f. energy is applied tothe tissue via the electrodes in a succession of pulse bursts ofprogressively increasing length and progressively decreasing amplitude.29. A method according to claim 28, wherein the increase in length ofsuccessive bursts is related to the decrease in amplitude of thosebursts.
 30. An electrosurgical system comprising the combination of anelectrosurgical generator as claimed in claim 1 and a bipolarelectrosurgical instrument coupled to an output of the generator.
 31. Asystem according to claim 30, wherein the instrument comprises a pair offorceps.
 32. A system according to claim 30, wherein the instrument isremovably connectible to the generator and includes an instrumentidentification element, and wherein the generator includes a sensingcircuit for sensing the identification element, the pulsing circuit ofthe generator being arranged automatically to adjust the pulse width ormark-to-space ratio if the said succession of pulses in response to theidentification element as sensed by the sensing circuit.
 33. A systemaccording to claim 32, comprising a plurality of bipolar electrosurgicalforceps instruments which are selectively connectible to the generatorand contain respective identification elements, wherein the electrodesof the instruments define different tissue contact areas, and whereinthe respective identification elements, the sensing circuit and/or thepulsing circuit and selected and configured to set the said pulse widthor mark-to-space ratio to a lower value for an instrument withelectrodes defining a comparatively large tissue contact area to ahigher value for an instrument with electrodes defining a comparativelysmall tissue contact area.
 34. A system according to claim 32, whereinthe identification elements, the sensing circuit and/or the pulsingcircuit are selected and configured to decrease the pulse frequency whenan instrument with a comparatively large tissue contact area isselected.
 35. An electrosurgical generator comprising a source of radiofrequency (r.f.) energy, at least a pair of output terminals forconnection to a bipolar electrosurgical instrument and for deliveringr.f. energy to the instrument, a pulsing circuit for the source, andcontrol circuitry including means for monitoring at least one electricalparameter associated with the output terminals, wherein the arrangementof the control circuitry, the pulsing circuit and the source is suchthat, with the output terminals connected to a resistive load, thecontrol circuitry causes the source to deliver into the load anamplitude-modulated r.f. power signal which, at least in an initialperiod, is at least one pulse with a predetermined ON-time, the or eachsuch pulse being followed by a predetermined OFF-time in whichsubstantially zero power is delivered and, in a subsequent period, has adifferent characteristic, the transition from the initial period to thesubsequent period being controlled by the control circuitry in responseto the at least one monitored parameter.
 36. A generator according toclaim 35, wherein the ON-time is less than or equal to 300 ms and theOFF-time is longer than the ON-time.
 37. A generator according to claim36, wherein the ON-time is between 40 ms and 300 ms.
 38. A generatoraccording to claim 36 or claim 37, wherein the OFF-time is between 100ms and 700 ms.
 39. A generator according to claim 35, wherein thecontrol circuitry is arranged to cause the r.f. power signal to provide,during the subsequent period, continuous energy delivery or more nearlycontinuous energy delivery than during the initial period.
 40. Agenerator according to claim 35, wherein the control circuitry isarranged such that the source delivers an r.f. power signal which has apeak power during the subsequent period which is lower than the peakpower in the initial period.